1. Nuclear Medicine Systems: Basics and Isotopes SPECT Instrumentation
BMME 560
Medical Imaging: X-ray, CT, and Nuclear Methods
Nuclear Medicine Systems: Basics and Isotopes
SPECT Instrumentation
Guest Lecturer:
Marijana Ivanovic
Office:
Radiology, 2112 Old Clinic
Tel:

2. Today Emission vs. Transmission imaging
Basic (“desired”) properties of radionuclides used for imaging
Detectors - configuration and characteristics
Planar Imaging
list mode, static, dynamic, gated, whole body
SPECT Instrumentation
Assignment 6

3. Emission vs. Transmission Imaging
X-ray methods : Transmission Imaging
Measure attenuation coefficient  ANATOMY
Detector
N=N0e-mx N0
Radiation position (direction) is known
Intensity of source is known (known flux (mAs) and energy (kVp))
X-ray tube

4. Emission vs. Transmission Imaging
Nuclear Medicine methods : Emission Imaging
Measure concentration and distribution of radiopharmaceutical in the body  PHYSIOLOGY (Organ Function, not structure)
Detector
Radiation position (direction) is NOT known
Intensity of source is NOT known
Energy is known

5. Nuclear Medicine Imaging – Basic Step:
Production of radionuclide
Labeling of radionuclide with pharmaceutical (tracer)
Injection (or inhalation) of radiopharmaceutical into patient
Wait for distribution and uptake of tracer in the organ of interest
Imaging (detect g-from radionuclide decay)
Tracer Injection
Uptake
Time: sec-day
Distribution

6. Requirements for Radiotraces
Modes of Radioactive Decay:
Gamma-ray emission (g)
- Isomeric Transition (IT)
- Internal Conversion (IC)
Alpha (a) emission
Beta minus (ß-) emission and (ß-, g)
Positron (ß+) emission and (ß+, g)
Electron capture (EC) and (EC, g)
Need g-ray emitters (exception:b+ emitters – ß+rapidly annihilated with electrons and produce g-ray).
Charged particles (a,b-) cannot penetrate tissue for emission imaging.

7. Requirements for Radiotraces
Decay scheme complexity:
Decay scheme complexity:
Desire isotopes relatively simple decay scheme – ideally one or two g-rays, no b- or a-rays. (b- or a-rays only increase radiation dose to the patient)
Very good imaging radiotracer 43
99m Tc
(6.01 h) g1
Not so good imaging radiotracer g3 g2
(140.5 keV - 89 %) 43 99 Tc
(1.2 • 105 y)

8. Requirements for Radiotraces
Energy of g-rays:
If the energy is “too low” a majority of the photons will be attenuated and will not reach the detector  simultaneously reduces the signal and adds radiation dose to the patient.
If the energy is “too high” a majority of the photons will pass through the detector without interacting in the detector (also difficult to collimate). 10 8 6 4 2 20 40 60 80
100
cm of water (muscle)
% transmitted
(511 keV)
99mTc (140 keV)
201Tl (80 keV)
g-ray energy should be high enough not to attenuate too much in the body, but low enough to be absorbed by the detector.
Energies of keV are used

9. Requirements for Radiotraces
Radiation decays exponentially and its characterized by a “half-life” T1/2 :
A(t) = A0 e- (ln 2* t /T 1/2)
Half-life:
If the half-life is “too short” it does not permit production, preparation (labeling) delivery, administration and internal distribution for imaging.
If the half-life is “too long” it will take to long to create an image and patient motion will be a problem. Radiation dose to the patient is increases with the half-life, due to a large number of radioactive atoms for a given activity .
Radiotracer preparation time
Imaging Time
Long T1/2
Radiation dose to the patient ~ with the area under the curve
Short T1/2
 Typical Half-lives are on the order of minutes to a few days.

10. Radionuclide Tp1/2 E (keV) Positron E(keV)
Physical Half-life (Tp1/2) and Photon Energy for Radionuclides commonly used in Nuclear Medicine
Radionuclide Tp1/ E (keV) Positron E(keV)
Technetium-99m (99mTc) 6.02 hr 140
Iodine-123 (123I) 13.3 hr 159
Indium-111 (111In) 2.82 d 173, 247
Thallium-201 (201Tl) 3.08 d 70, 167
Gallium-67 (67Ga) 3.25 d 92, 184, 296
Xenon-133 (133Xe) 5.31 d 81
Iodine-131 (131I) 8.05 d 364
Iodine-125 (125I) 60.2 d 35, 27
Fluorine-18 (18F) m
Carbon-11 (11C) 20.3 m
Nitrogen-13 (13N) 10.0 m
Oxygen-15 (15O) 2.1 m
Rubidium-82 (82Rb) 1.3 m
In order to be useful the radionuclide must be “safe” and able to “trace” within the body, either by itself or attached to a compound.

11. 2-[18F] Fluoro-2-Deoxy-D-Glucose (FDG)
Requirements for Radiotraces
Chemical properties:
Must be able to incorporate isotope into a pharmaceutical or other organic compound.
2-[18F] Fluoro-2-Deoxy-D-Glucose (FDG)
The pharmaceutical part of the radiopharmacetical determines the biodistribution and organ uptake and clearance.

12. Nuclear Medicine Imaging:
Depending on the radiopharmaceutical (radiotracer), different physiological or biochemical functions are being imaged.
Cardiac imaging
99mTc-Sestamibi
Tumor imaging
18F-FDG
Bone Imaging
99mTc-MDP

13. Half–times : Physical, Biological, Effective
The amount of activity present in an organ after the injection generally changes with time, owing to physical decay of radionuclide and biological uptake and excretion processes.
M = A e-lb t
= (A0 e-lp t) e-lb t = A0 e-(lb+ lp) t = A0 e-leff t
T1/2b T1/2 p
T1/2 eff =
T1/2 b + T1/2p
T1/2 eff ≤ shorter of the two, T1/2p and T1/2b
   when T1/2p >> T1/2b, then T1/2eff ≈ T1/2b
(tracer excretes very fast – example: 99mTc DTPA, 133Xe)
  when T1/2b >>T1/2p, then T1/2eff ≈ T1/2p
(tracer does not excrete or excretes very slowly – examples: 201Tl, 99mTc MAA)
Imaging parameters and the amount of activity that could be injected depend on the effective half-life.

14. Basic Radiation Detector System for Nuclear Medicine Imaging
What do we need to know about the radiation?
Energy?
Position?
How much?
Detector
Signal
Signal Processing
(energy, position..)
Stored to disk
Incoming g-ray
What are the important properties of the detector?
Energy resolution
Spatial resolution
Sensitivity
Counting rate

15. Types of Detectors According to the type of information produced:
Counters - indicate the number of interactions that occur in the detector
Spectrometers – yield the information about the energy distribution of the incident radiation
Dosimeters – indicate the net amount of energy deposited in the detector by multiple interactions
Detector for NM imaging have to be Counters and Spectrometers.

16. Types of Detectors Scintillation detectors
Inorganic scintillators (NaI(Tl))
Scintillators are materials that emit visible or UV light following the ionization or excitation.
Solid State (semiconductor) detectors
Most of the detectors on clinical NM imaging systems are inorganic scintillators.
There are only few cameras with solid state detectors (small field of view),

17. Properties of Some Scintillator Materials
1.56
1.9
1.82
2.15
1.48
1.85
Index of refraction
Little No
Very
Yes
Hygroscopic
2250‡
3100#
4300
4200
4800
3900
4150
l of max. emission ( Å)
2.0
6.4 30
4.8
2.5 40
Photo yield** (per keV)
0.8 56
300
230
Scint. decay time (msec)* 60 59 66 74 53 50
Effective Atomic No. -
0.674
0.833
0.955
0.34
Attenuation coefficient
511 keV, cm-1)
4.89
6.71
7.4
7.13
4.61
3.67
r (g/cm3)
BaF2
GSO3
LSO2
BGO1
CsF
NaI(Tl)
Property
1Bi4Ge3O12 ; 2Lu2SiO5 ; 3Ge2SiO5
*Time required for emission of ~67% of the light
** Average number of scintillation photons emitted per keV of ionizing radiation energy absorbed.
‡ Fast component; # Slow component

18. Photo Multiplier Tube (PMT)
Scintillation detector consists of a scintillator and a light detector (photomultiplier tube, PMT).
Output signal
Input
window
Photocathode
Light
Photon
400 V
300 V
500 V
1200 V C 1 3 9 27 81
60000
. . . e
Focusing
grid
dynode
anode
High
Voltage
PMTs are electronic tubes that produce a pulse of electrical current when stimulated by weak light signal.
Total electron multiplication is very large : ~ 610 ( ~6x107) for 10 stage dynode.

19. Inorganic Scintillation Detectors (Scintillator+PMT)
Detectors on majority clinical NM planar and SPECT imaging systems use NaI(Tl) crystals.
NaI - in pure state is scintillator at liquid nitrogen temperatures
NaI(Tl) - scintillator at room temperatures
g-photo
Input window
Scintillation center
NaI(Tl) crystal and PMT assemblies
MgO ili Al2O3
reflector
NaI(Tl)
Visible
Light
Photons
Aluminum or
Steel shield
Glass
window
PMT
mmetal shield

20. Advantages of NaI(Tl) detectors:
Detectors on majority clinical NM planar and SPECT imaging systems use NaI(Tl) crystals. NaI(Tl) is excellent for single-photon detectors.
Advantages of NaI(Tl) detectors:
It is relatively dense (r=3.67 g/ cm3) and contains an element of relatively high atomic number (iodine, Z=53). Therefore it is a good absorber and efficient detector of penetrating radiations, such as x- rays and g-rays.
It is a relatively efficient scintillator, yielding one visible light photon per approximately 30 eV of radiation energy absorbed.
It is transparent to its own scintillation emissions. Therefore there is little loss of scintillation light caused by self-absorption, even in NaI(Tl) crystal of relatively large size.
A NaI(Tl) detector provides an output signal (from PM tube) that is proportional in amplitude to the amount of radiation energy absorbed in the crystal. Therefore it can be used for energy selective counting.

21. Disadvantages of NaI(Tl) detectors:
The NaI(Tl) crystal is quite fragile and easily fractured by mechanical or thermal stresses (e.g., rapid temperature changes). Fractures in the crystal do not necessarily destroy its usefulness as a detector, but they create opacifications within the crystal that reduce the amount of scintillation light reaching the photocathode.
Sodium iodide is hygroscopic. Exposure to moisture or a humid atmosphere causes a yellowish surface discoloration that again impairs transmission to the PMT. Thus hermetic sealing is required.
Sodium iodide crystals of large size (30-50 cm diam) are difficult to grow and quite expensive.

22. Semiconductor detectors
Solid-state analogs of gas-filled ionization chambers.
When ionizing radiation interacts with the detector, electrons in crystal are raised to an exited state, permitting an electrical current to flow.
times more dense the gas --> much better stopping power and more efficient for x- and g- rays ( one ionization per 3 eV).
Usually requires very high purity materials or introduction of “compensating” impurities that donate electrons to fill electron traps caused by other impurities.
Count individual events
Size of electrical signal is proportional to the energy absorbed
Advantages: superb energy resolution
Disadvantages: - high "noise current at room temperature - have to be cooled at 77° K (-196°C)
Limited crystal size ( 5x5 cm) and very expensive

23. “New” semiconductor detectors
CdTe and CZT are less well-developed semiconductor materials that overcome two of the major disadvantages of Si and Ge:
they can be operated on room temperatures without excessive electronic noise
their high atomic number means that a relatively thin detectors have a good stopping power for detecting g rays.
Although CdTe and CZT are now being used in some nuclear imaging counting and imaging devices, their use has been restricted to smaller detectors because of difficulty and expense of growing large CdTe and CZT with required purity.

24. Interactions of photons with a spectrometer:
A - Photoelectric
B - Compton+Photoelectric
C - Compton
D - Photoelectric with characteristic x-ray escape
E - Compton scattering in the shielding and scattered photons enters the detector
F - Characteristic x-ray from lead shield
NaI(Tl)
PMT
Lead Shield F E A B C D
Source

25. Sample Spectra with Cs-137
Actual energy spectrum
Spectrum obtained with a NaI(Tl) detector
662 keV
32 keV
662 keV
90% keV g-ray
Number of interactions
Number of interactions
32 keV
10% electron conversion
Followed by a ~32 keV
K-shell x-ray
200
400
600
800
200
400
600
800
Energy (keV)
Energy (keV)
Due to different (partial deposition) of the energy.
Statistical fluctuations in the process by which the energy is deposited and converted into an electrical signal (random variations in a fraction of deposited energy converted to light, fraction of light that reaches PMT and number of electrons ejected from photocathode per unit energy deposite by the light,…)

26. Acquisition & processing computer
Scintillation Camera (Gamma camera)
Collimator
NaI Crystal
PMT
Lead Shield
Source
Electronic boards
Acquisition & processing computer

27. lead shield
collimator

28. Gamma camera

29. Collimators Image of the source No With Collimator Collimator Source
To obtain image with the gamma camera, it is necessary to project g-rays from the source distribution onto the camera detector.
Gamma rays can not be focused, therefore most practical way to project g-rays on an imaging system employs the principle of absorptive collimation for image formation. An absorptive collimation projects an image of the source distribution onto the detector by allowing only those g-rays traveling along certain direction to reach the detector. g-rays not traveling in proper direction are absorbed by the collimator before they reach the detector.
“Projection by absorption” technique is very inefficient method for utilizing radiation because most of the potentially useful radiation traveling toward the detector is actually stooped by the collimator.
Image of
the source No
Collimator
With
Collimator
Source

30. Collimator system is the heart of imaging system – it has the biggest impact on SNR.
Its function is to form an image by determining the direction along which gamma-ray propagates. L d
L/2
d/2
Collimator’s Resolution and Sensitivity are determined by the ratio of a collimator hole diameter (d) and length (L).
L  (d=const. )  Resolution  & Sensitivity 
d  (L=const. )  Resolution  & Sensitivity 

31. NaI(Tl) Crystal 3/8”(up to 1”) thick
Density = 3.67 g/cm3
Attenuation 140 keV = cm-1  3/8” stops 92% photons
PE fraction = 80%
Scint. decay time = 230 nsec
Photon yield - 40/keV  40*140 = 5,600 light photons emitted for each detected photon
Resolution & Efficiency vs. Crystal Thickness
Crystal
Thickness
(Inches)
FWHM
(mm)
Photopeak
efficiency
@140 keV
@ 511 keV
1/4
3.0
0.70 -
3/8
3.5
0.80
0.055
1/2
3.7
0.85
0.07
5/8
3.9
0.90
0.09
3/4
4.4
0.96
0.10
1.0
4.5
0.99
0.30
Thinner the crystal better resolution, but lower efficiency.

32. Spatial Positioning Z X k = Z Y k =
SMC X- X+ Y- Y+
PMT’s
NaI(Tl)
crystal Z X k - + = Z Y k - + =
The summing matrix circuits (SMC) combine the signals from the individual PM tubes in such a way that the relative amplitudes of the X+ and X- signals, and of the Y+ and Y- signals, are proportional to the distance of the scintillation event from the center line of the crystal. These four signals, are used to determine the location at which the scintillation event occurred, and the Z signal can be used to determine the energy. Z signal (combined output of from all PM tubes) is proportional in amplitude to the total amount of light produced by a scintillation event in the crystal.

33. Example of light distribution over PMT locations for 37 tube camera. 9 11 6 3 1 2 1
voltage 78 74 82
424 1
FOV 1 2 3 6 11 5 12 73 67 8
423 13 76
Example of light distribution over
PMT locations for 37 tube camera.
The scintillations are centered
over the highlighted locations.

34. Gamma Camera Energy Spectra (Summed signal from all PMTs)20 40 60
100 80 50
150
200
Energy (keV)
Pulse height (a.u.)
Source in air
Source in water
Scattered photons
Scattered photons are mis-positioned in the image (reduce image contrast)
Energy Windows
Balance between accepting all good events and rejecting scattered photons.
Most camera can acquire multiple (4-8) energy windows simultaneously
Energy resolution of new generation of gamma-cameras is 8-10%.

35. Image Acquisition Frame Mode acquisitions: Static Dynamic Gated
single or mutiple images acquired at different times and/or different angles
can have multiple energy windows
Dynamic
series of images acquired sequentially
Gated
repetitive, dynamic imaging (used for cardiac imaging)
Whole Body
Continuous or “step&shoot” table motion during acquisition

36. Static Acquisition
Different views of the same organ AP PA

37. Dynamic acquisition
Time
20 min
1 frame /min

38. Dynamic Acquisition - Processing
Time
Time 1 3 5 7 9 11 13 15 17 19
L. Kidney
ROI counts
Bladder
R. Kidney
aorta
Time

39. Whole Body Bone Scan
Multiple foci not corresponding to traumatic or articular pattern

40. Gated Acquisition ECG Each Image is contribution from 600 heart cyclesP T P T P T 1 16
Time
Each Image is contribution from 600 heart cycles
Left Ventricle Curve
Ejection Fraction 1 16

41. Single Photon Emission Computed Tomography
SPECT
Single Photon Emission Computed Tomography
Projections
Reconstructed
Transaxial Slices

42. To increase efficiency: Most SPECT cameras have several detectors
From: IEEE TNS VOL. 42, NO. 4, 1995,Jingai Liu, Wei Chang, and Srecko Loncaric
Use fan-beam or cone-beam collimation when imaging small FOV (brain, pediatrics, cardiac..)
Dedicated systems for Brain and Cardiac imaging are considered.

43. Dedicated Brain SPECT systems
inSPira HD portable SPECT scanner for brain imaging from NeuroLogica corporation
inSPira HD features spiral-rotating focused collimators. Image quality approaches PET with the resulting reconstructed spatial resolution as high as 3.0mm.
The focused collimators and spiral scan motion of the inSPira HD are responsible for the higher resolution (both in-plane and in z-axis) as compared to conventional Gamma Camera SPECT systems. The combination allows isotropic scans and reconstruction with as high as 3mm resolution in X,Y,Z.

44. What is needed for SPECT?
Complete set of projections for each axial plane.
The radiotracer distribution must is stable
The detector is always viewing the same distribution
Patient is not move during the acquisition
Imaging system must be in alignment and have uniform and stable detectors
Assumptions:
SPECT Imaging NO
YES
Activity in Organ
Time

45. Artifacts due to incomplete angular sampling:
Rods
Spheres
Uniform
Complete set
of projections (360˚)
Incomplete set
of projections (240˚)
Missing projections

46. 180° acquisition is allowed only for Cardiac SPECT
due to anatomical position of the heart
Reconstructed region
Acquiring over 360 ° reduces some of the inconsistency associated with SPECT and reduces distortion, BUT
Tissue attenuation degrades the quality of the projections collected from the posterior
The Nuclear Cardiology community has overwhelmingly endorsed 180 ° orbits

47. Effect of Collimator Damage (Nonuniformity) on Reconstructed Images
Flood image of Head 2
before damage
Flood image of Head 2
after damage
Damaged Collimator
on Head 2
Integral Unif. = 2.5 %
Differential Unif. = 1.9 %
Integral Unif = 4.2 %
Differential Unif. = 3.4 %
SPECT study
before damage
SPECT study
after damage

48. Quantitative SPECT? Collimator Blur Attenuation Scatter
The ultimate goal of quantitative SPECT is to provide reconstructed
images in which each pixel value in the image represents the absolute activity concentration in the corresponding region in the patient.
The most important factors affecting quantitation are:
Collimator Blur
Attenuation
Scatter

49. Collimator Blur:
Source to Collimator distance vary as camera rotates during SPECT imaging.
Tangential
Gamma
Camera
Radial
FBP
Different Radial and
Tangential Resolution

50. Depth Dependant Resolution Recovery
From : M. O’Connor MWSNM-04 presentation

51. Attenuation: 25.0% 50.0 % 6.25 % 37 cm 12.5 % 28 cm 38 cm 41 cm
Average Size Patient
Large Patient

52. No Attenuation Correction:
Attenuation Correction Applied:

53. Accurate Attenuation Correction
Iterative Algorithm
Filter backprojection cannot incorporate attenuation correction
Co-registered Attenuation Map
Transmission measurements
Transmission reconstruction
Attenuation coefficient conversion

54. Emission + Transmission SPECT
Attenuation Compensation
Emission + Transmission SPECT
Gantry mounted x-ray tube emits x-rays to an opposing CT detector (GE)
Multiple line sources in “wings” (Siemens)
Multiple Point Sources (Philips)
Scanning line sources (Adac)

55. From : M. O’Connor MWSNM-04 presentation

56. From : M. O’Connor MWSNM-04 presentation

57. Attenuation Correction using Transmission Scan
Uncorrected
Corrected

65. Scatter:
In a typical patient study with 99mTc labeled radiopharmaceutical, even using narrow 15% PHA window, the ratio of the number of detected scattered photons to the number of nonscattered photons may be as large as 40%.
The presence of scattered photons results in reduced image contrast and leads to an overestimation of the concentration of radioactivity in the pixel. The loss of image contrast may obscure clinically important details , particularly "cold" areas in the images.
20 cm dia. cylinder filled with 99mTc and 6 cm dia. cold sphere
180
140
100 60 20
Energy (keV)
Counts (arb. units)
15% window
Scattered photons
Measured line profile
"Ideal" line profile

66. Scatter Correction Methods:

67. http:// http://www.physics.usyd.edu.au/ugrad/sphys/medphys.html

70. New dedicated SPECT systems:
Detector Specifications:
Detector type: pixilated (16x64)
Detection material: CdZnTe
Number of detectors: 9
Field of view: 15.74cm x 3.94cm
Collimator: built-in tungsten array
Energy resolution: <8% (99mTc)
System uniformity: Integral: <4%
Differential: <3%
System Planar Sensitivity: 500 [cnts/μCi/min]
Tungsten + Lead Shielding: 170keV
Detailed diagram of a single-detector column from D-SPECT camera.

71. D-SPECT™
Conventional Camera:
16 min.
20 min.
D-SPECT:
2 min
4 min

72. CardiArc
Figure 6. Imaging with the CardiArc camera is accomplished via 3 curved NaI(Tl) crystals and an array of photomultiplier tubes. Collimation is achieved via a thin lead sheet with 6 vertical slits (aperture arc), which rotates during acquisition. (Courtesy of Dr. Jack Juni of CardiArc.)
Figure 7. Slice separation is accomplished with the CardiArc camera by use of thin lead vanes that are stacked vertically to define slices for imaging. (Courtesy of Dr. Jack Juni of CardiArc.)
From: Journal of Nuclear Cardiology, Patton, Slomka, and Berman, Volume 14, Number 4; Recent technologic advances in nuclear cardiology

74. Cardius 3 XPO triple-head, pixilated detector camera (Digirad).
Figure 3. Data acquisition with Cardius 3 XPO camera. The detectors remain fixed while the patient is rotated through 202.5° via a rotating chair configuration
Each detector head is cm and contains an array of –mm thick CsI(Tl) crystals coupled to individual silicon photodiodes used to convert the light output of the crystals to electrical pulses. Digital Anger logic is used to process the signals and create images. In the 3-detector system the detector heads are fixed in position at 67.5° between heads. For imaging, the patient sits on a chair with his or her arms placed on an arm rest above the detectors. With this system, the manufacturer reports a reconstructed spatial resolution of 15.4 mm and a sensitivity of 234 cpm/mCi using the system’s cardiac collimator.

75. GE Healthcare - Alcyone Technology, a nuclear cardiology platform combining cadmium zinc telluride (CZT) detectors, focused pin-hole collimation, 3D reconstruction, and stationary data acquisition, to improve workflow, dose management, and overall image quality.